1. Field of the Invention
The present invention relates to a method and apparatus employing detector pixels for obtaining an image having a resolution which is not directly related to the sizes of the detector pixels. More particularly, the present invention relates to a method and apparatus which obtains a series of spatially filtered high-resolution digital x-ray or gamma ray images of portions of an object or objects while minimizing image degradation due to conversion blurring and radiation scattering, and which arranges the spatially modulated images into a larger complete image of the object or objects.
2. Description of the Related Art
Various techniques currently exist and many are under development for obtaining digital x-ray and gamma ray images of an object for purposes such as x-ray diagnostics, medical radiology, non-destructive testing, and so on. Known devices include line digital detectors, which obtain images along essentially one direction, and therefore must be scanned across an object to obtain sectional images of the object which can be arranged into an image of the entire object. Also known are two-dimensional digital detectors which can obtain an image of the entire object at one time, and thus can operate faster than an apparatus which includes a line detector.
A digital x-ray imager creates a digital image by converting received x-rays, which are used to form the image, into electrical charges, and displaying the charge as a function of position. Digital x-ray detectors typically have the potential of high sensitivity and large dynamic range. Therefore, when used in medical applications, a digital x-ray detector will generally be capable of obtaining a suitable image of the patient without requiring the patient to receive a large dose of x-ray radiation.
Digital image data is also much easier to store, retrieve and transmit over communication networks, and is better suited for computer-aided diagnostics, than conventional film x-rays. Digital x-ray images can also be displayed more easily than conventional film x-rays, and provide greater image enhancement capabilities, a faster data acquisition rate, and simplified data archival over conventional film x-rays. These advantages make digital x-ray imaging apparatus more desirable than film x-ray apparatus for use in many diagnostic radiology applications, such as mammography.
The general construction and operation of digital x-ray detectors will now be described. As discussed briefly above, digital x-ray detectors collect electrical charges produced by x-rays as a function of position, where the amount of charge is directly proportional to the x-ray intensity. Two general approaches for x-ray conversion are currently under investigation for flat-panel digital x-ray detectors. These approaches are generally referred to as the indirect method and the direct method.
In the indirect method, x-rays are converted to low-energy photons by a scintillator, and the low-energy photons are then converted to electrical charges by solid-state detectors. This method is described in a publication by L. E. Antonuk et al., "Signal, Noise, and Readout Considerations in the Development of Amorphous Silicon Photodiode Arrays for Radiotheraphy and Diagnostic Imaging" Proc. SPIE 1443:108 (1991), the entire contents of which is incorporated by reference herein.
In the direct method, x-rays are converted to electron-hole pairs by photoconductors. An electric field applied to the photoconductor separates the electrons from the holes. This method is described in a publication by J. A. Rowlands et al. entitled "Flat Panel Detector for Digital Radiology Using Active Matrix Readout of Amorphous Selenium," Physics of Medical Imaging SPIE 3032: 97-108(1997), and in an article by R. Street, K. Shah, S. Ready, R. Apte, P. Bennett, M. Klugerman and Y. Dmitriyev, entitled "Large Area X-Ray Image Sensing Using a PbI.sub.hd 2 Photoconductor," Proc. SPIE 3336: 24-32 (1998). The entire contents of both of these papers are incorporated by reference herein. Many types of photoconductors are under development by medical imaging community.
A type of flat-panel, two-dimensional, digital x-ray, imager comprises a plurality of charge-coupled devices (CCDs) on a silicon substrate. The CCDs can be easily made on the silicon substrate to have a pixel pitch smaller than 10 .mu.m.times.10 .mu.m. However, because the maximum size of silicon substrates is limited, to achieve the dimensions needed for a large-area flat-panel x-ray detector, multiple wafers have to be patched together. Some of the CCD x-ray detectors are described in the following publications: F. Takasashi, et al., "Development of a High Definition Real-Time Digital Radiography System Using a 4 Million Pixels CCD Camera", Physics of Medical Imaging SPIE 3032: 364-375 (1997); J. M. Henry, Martin J. Yaffe and T. O. Tumer, "Noise in Hybrid Photodiode Array--CCD X-ray Image Detectors for Digital Mammography," Proc. SPIE 2708: 106 (1996); and M. P. Andre, B. A. Spivey, J. Tran, P. J. Martin and C. M. Kimme-Smith, "Small-Field Image-Stitching Approach to Full-View Digital Mammography," Radiology 193, Suppl. Nov.-Dec., 253-253 (1994), the entire contents of each being incorporated by reference herein.
Alternatively, a flat-panel imager can include active matrix arrays of thin film transistors (TFTs) on a glass substrate. Because glass substrates can be large, the digital x-ray imager can, in principle, be made of a single substrate. However, it is very difficult to make a digital detector with a pixel pitch much smaller than 100 .mu.m using substrates other than silicon wafers, as described in the following publications: L. E. Antonuk et al., "Development of Thin-Film, Flat-Panel Arrays for Diagnostic and Radiotherapy Imaging", Proc. SPIE 1651: 94 (1992); L. E. Antonuk et al., "Large Area, Flat-Panel, Amorphous Silicon Imagers", Proc SPIE 2432: 216 (1995); and L. E. Antonuk et al., "A Large-Area, 97 .mu.m Pitch, Indirect-Detection, Active Matrix Flat-Panel Imager (AMFPI)", SPIE Medical Imaging 1998 Technical Abstracts, San Diego, 83 (1998), the entire contents of each being incorporated by reference herein.
As discussed above, digital x-ray imaging techniques represent a vast improvement over conventional film x-ray apparatus. However, digital x-ray imaging systems experience certain drawbacks with regard to image resolution.
It has been a common belief that the resolution of the digital image can be no better than the pixel pitch (pixel periodicity) of the imaging apparatus, and is rather often much worse due to various types of blurring phenomena which occur during image acquisition. However, as can be appreciated from the description of the operation of digital x-ray detectors set forth below, pixel pitch is only one of the many factors that influence the resolution of a digital image obtainable by a digital imaging apparatus.
Detectors for digital radiography are composed of discrete pixels which generally have a uniform size, shape and spacing. The "fill factor" is defined as the active portion of each detector pixel that is used for charge collection relative to pixel pitch or, in other words, the fraction of the pixel area occupied by the sensor for x-ray detection. A flat-panel imager having thin-film transistors (TFTs), for example, has a fill factor which decreases dramatically as the pixel pitch decreases. The TFTs are large compared to transistors on silicon substrates, and the various electrode lines occupy much surface area of the glass substrate. Hence, the fill factor decreases greatly as the pixel pitch decreases.
For example, the fill factor is 57% for a 127 .mu.m pixel pitch array, and is 45% for a 97 .mu.m pixel pitch array which performs indirect x-ray conversion and has been aggressively designed, as described in the article entitled "A Large-Area, 97 .mu.m Pitch, Indirect-Detection, Active Matrix Flat-Panel Imager (AMFPI)" cited above.
The fill factor approaches zero as the pixel pitch decreases toward 50 .mu.m in a detector employing indirect converters. When the fill factor is small, the sensitivity of the detector suffers greatly. Fortunately, however, the fill factor can be improved using direct x-ray converters and a vertical stacking architecture. However, such device becomes increasingly difficult to fabricate as pixel pitch decreases. Thus, development costs for such a device are very high, and it is unclear what the smallest achievable pixel pitch could be with this technique.
In addition, connecting the data and control lines from the detectors to the gate driver chips and readout amplifiers of the pixel array presents severe packaging problems. Currently, bonding of large array of leads from substrate to cable is limited to a device having no less than about an 80-100 .mu.m pixel pitch. By increasing the pixel resolution, multiplexed contacts or new bonding techniques must be developed to create input and output terminals for the device.
The modulation transfer function (MTF), which is a function of spatial frequency f versus location on the detector, is useful for analyzing spatial resolution. Larger MTF values mean better resolution. For existing flat-panel detectors, MTFs are important in analyzing two steps of the image acquisition sequence: the detector pixel pitch, and the blurring produced during the conversion of x-rays to charges. (See, e.g., an article by J. M. Henry, Martin J. Yaffe and T. O. Tumer, "Noise in Hybrid Photodiode Array--CCD X-ray Image Detectors for Digital Mammography," Proc. SPIE 2708:106(1996), the entire contents of which is incorporated by reference herein).
The charges generated by x-ray conversion can become blurred spatially. The source of blurring for indirect conversion using phosphor is different from that for direct conversion. For most detectors, the measured MTF is dominated primarily by the blurring of the converter when the pixel pitch is 100 .mu.m or smaller.
In addition, settled phosphor scatters light generated by the x-rays. The lateral spreading of the light is approximately equal to the thickness of the layer. For settled phosphor, spatial resolution becomes finer, but the quantum efficiency decreases as the thickness of the phosphor decreases. Optimized thin photoconductors are expected to produce smaller spread. Although the light spread may be less of a problem for thick collimated CsI phosphor, the boundaries of the CsI grains are not perfect.
Furthermore, spatial resolution can be degraded due to x-rays striking the detector at an oblique angle. This problem exists for both direct and indirect x-ray converters. The extent of the charge spread collected by the detector is a function of the incidence angle. Since the x-ray incidence angle is a function of location on the detector relative to the x-ray point source, the modulation transfer function (MTF) of conversion blurring and oblique x-ray incidence blurring MTF.sub.conversion is also a function of the location on the detector. The MTF.sub.conversion of for Lanex Regular is much worse than Lanex Thin. The MTF of for Lanex Thin is 0.2, at 5 cycles/mm, as described in the article entitled "A Large-Area, 97 .mu.m Pitch, Indirect-Detection, Active Matrix Flat-Panel Imager (AMFPI)", cited above.
The final system MTF is the product of the MTF associated with various components of the system, including the detector array MTF introduced by the detector pixel pitch and the MTF of conversion blurring. For these reasons, the reduction of pixel pitch alone is not as good the combination of reduction of pixel pitch and reduction of conversion blurring. The resolution of the detector is also effected by a variety of other factors that will not be discussed in detail here such as signal statistical noise, charge conversion noise and electronic noise.
Gamma rays are radiation generated by nuclear process. The energy of gamma rays are typically higher than that of the x-rays, but low energy range of the gamma rays can overlap the high energy end of the x-rays. These detector concepts can also be applied to the detection of gamma rays and megavolt radiation. A thick scintillator or a metal plate/phosphor screen combination is used. This is described in a publication by L. E. Antonuk, et al., "Demonstration of Megavoltage and Diagnostic X-ray Imaging with Hydrogenated Amorphous Silicon Arrays," Med. Phys. 19: 1455 (1992), the entire contents of which is incorporated by reference herein.
In summary, the major problems expected with small pixel detector development are complicated circuit architecture, increased number of leads to be bonded, the small pitch of the leads necessary for bonding, and resolution being increasingly dominated by scintillator blurring and the oblique x-ray incidence effect. These drawbacks result in decreased manufacturing yield, high risk and expensive development.
Accordingly, a continuing need exists for an apparatus capable of obtaining high-resolution digital x-ray or gamma ray images without the drawbacks discussed above.